1. Field of the Invention
The present invention generally relates to the use of sensors to monitor the concentration of a chemical species in bodily fluids.
2. Description of the Related Art
There exist numerous medical conditions which require monitoring of a chemical species within bodily fluids of a patient. For example, diabetes mellitus is a serious medical condition affecting approximately 10.5 million Americans, in which the patient is not able to maintain blood glucose levels within the normal range (normoglycemia). Approximately 10% of these patients have insulin-dependent diabetes mellitus (Type I diabetes, IDDM), and the remaining 90% have non-insulin-dependent diabetes mellitus (Type II diabetes, NIDDM). The long-term consequences of diabetes include increased risk of heart disease, blindness, end-stage renal disease, and non-healing ulcers in the extremities. The economic impact of diabetes to society has been estimated by the American Diabetes Association at approximately $45.2 billion annually (Jonsson, B., The Economic Impact of Diabetes, Diabetes Care 21(Suppl 3): C7-C10, (1998)).
A major long-term clinical study, the Diabetes Control and Complications Trial, involving 1,441 patients with insulin-dependent diabetes mellitus (Type I diabetes) over a 10-year period from 1984-1993, demonstrated that by intensive therapy (frequent administration of either short- or long-acting insulin), these long-term consequences (retinopathy, nephropathy, and neuropathy) could be reduced (“The Effect of Intensive Treatment of Diabetes on the Development and Progression of Long-Term Complications in Insulin-Dependent Diabetes Mellitus,” The Diabetes Control and Complications Trial Research Group, New Eng. J. Med., 329: 977-86 (1993)). Unfortunately, a major difficulty encountered during the trial was that intensive treatment also resulted in a higher incidence of low blood glucose levels (hypoglycemia), which was severe enough to result in coma or death, as compared to patients under conventional medical management.
Currently, diabetics must monitor their condition by repeatedly pricking their fingers in order to obtain blood samples for evaluation. The major drawback to self-monitoring of glucose is that it is discontinuous and therefore the number of glucose measurements performed is dependent on the motivation of the patient.
Existing analytical techniques and devices for in vitro glucose measurements have a high level of accuracy (the error can be <1%). Many of these routine methods are accepted as standards of comparison with new devices. Management of diabetes currently relies on these methods to control the disease and minimize complications.
There are two main disadvantages to these existing options. First, sampling even a minimal amount of blood multiple times per day is associated with risks of infection, nerve and tissue damage, and discomfort to the patients. Second, in the case of dynamic changes in glucose concentration, very frequent or even continuous measurements of blood glucose levels are required (Wilkins, E., et al., “Glucose Monitoring: State of the Art and Future Possibilities”, Med. Eng. Phys. 18(4):273-88, (1996).
There are two main approaches to the development of a continuous blood glucose monitor. The first category is non-invasive sensors, which obtain information from physico-chemical characteristics of glucose (spectral, optical, thermal, electromagnetic, or other). The second category is invasive sensors. In this group, there is intimate mechanical contact of the sensor with biological tissues or fluids, because the device is placed within the body. (Wilkins, 1996).
Non-invasive sensor technology has focused on the absorption of the near-infrared (NIR) spectra by the analyte, in this case, glucose (See U.S. Pat. No. 5,945,676 to Khalil, et al., and U.S. Pat. No. 5,433,197 to Stark). Absorptions which occur in the NIR region are most often associated with overtone and combination bands of the fundamental vibrations of —OH, —NH, and —CH functional groups. As a result, most biochemical species will exhibit some absorption in the region of interest. Glucose measurements are usually performed in the spectra region from 4250 to 660 cm−1. These highly overlapping, weakly absorbing bands were initially thought to be too complex for interpretation and too weak for practical application. Improvements in instrumentation and advances in multivariate chemometric data analysis techniques may allow meaningful results to be obtained from these complex spectra.
However, to date these devices are not particularly accurate even in the normal physiological range. A subject-dependent concentration bias has been reported. The temperature sensitivity of water absorption bands in the glucose-measuring region can be a significant source of error in clinical assays. In addition, the devices can also be affected by individual variations between patients at the measurement site. Skin location, temperature and tissue structure may affect the results, and decrease the accuracy of the reading.
Other investigators have looked into measurement of glucose from body fluids other than blood, such as sweat, saliva, urine, or tears. However, factors relating to diet and exercise can affect glucose levels in these fluids. In general, there is no strong correlation established between glucose concentration in the blood and in excreted fluids. The lag time between blood and excreted fluid glucose concentrations can be large enough to render such measurements inaccurate.
The continuous in vivo monitoring of glucose in diabetic subjects should greatly improve the treatment and management of diabetes by reducing the onus on the patient to perform frequent glucose measurements. Implanted glucose sensors could be used to provide information on continuously changing glucose levels in the patient, enabling swift and appropriate action to be taken. In addition, daily glucose concentration measurements could be evaluated by a physician. An implantable sensor could also provide an alarm for hypoglycemia, for example, overnight, which is a particular need for diabetics. Failure to respond can result in loss of consciousness and in extreme cases convulsive seizures. Similarly, a hyperglycemic alarm would provide an early warning of elevated blood glucose levels, thus allowing the patient to check blood or urine for ketone bodies, and to avert further metabolic complications. (Jaffari, S. A. et al., “Recent Advances In Amperometric Glucose Biosensors For In Vivo Monitoring”, Physiol. Meas. 16:1-15 (1995)).
Invasive glucose sensors may be categorized based on the physical principle of the transducer being incorporated. Current transducer technology includes electrochemical, piezoelectric, thermoelectric, acoustic, and optical transducers.
In piezoelectric, thermoelectric, and acoustic (surface acoustic wave, SAW) sensors used for glucose measurement, an enzyme-catalyzed reaction is used to create a measurable change in a physical parameter detected by the transducer. The development of these sensors is at an early laboratory stage (Hall, E., Biosensors, Oxford University Press. Oxford, 1990). Optical sensors are based on changes in some optical parameter due to enzyme reactions or antibody-antigen reactions at the transducer interface. Based on the nature of the monitoring process, they are densitometric, refractometric, or colorimetric devices. At present, none of them meets the selectivity requirements to sense and accurately measure glucose in real physiological fluids.
There is a significant body of literature regarding the development of electrochemical glucose sensors. These generally incorporate an enzyme, which selectively reacts with glucose. Examples of enzymes, which selectively react with glucose, are glucose oxidase (GOD), hexokinase, glucose-6-phosphate dehydrogenase (G-6-PD), or glucose dehydrogenase. Hexokinase is an enzyme that catalyzes the phosphorylation of glucose by ATP to form glucose-6-phosphate and ADP.

Monitoring the reaction requires a second enzyme, glucose-6-phosphate dehydrogenase, in the following reaction:

The formation of NADPH may be measured by absorbance at 340 nm or by fluorescence at 456 nm (Jaffari, 1995).
Glucose dehydrogenase is another enzyme, which may be used for monitoring glucose in the following reaction:
The NADH generated is proportional to the glucose concentration.
Glucose oxidase is the most commonly used enzyme reported in the literature. Its reaction is relatively simple, inexpensive, and may be monitored using a variety of techniques.
These advantages have led to the extensive use of this enzyme in clinical analysis as well as its incorporation in the majority of prototype biosensor configurations. The reaction of glucose with this enzyme is a two-stage reaction:-D-glucose+GOD(FAD)→glucono-δ-lactone+GOD(FADH2)  1)GOD(FADH2)+O2→GOD(FAD)+H2O2  2)glucono-δ-lactone+H2O→gluconic acid  3)The overall reaction is usually expressed as:β-D-glucose+O2+H2O gluconic→acid+H2O2  4)The reaction can therefore be monitored by the consumption of oxygen, the production of hydrogen peroxide, or the change in acidity due to the increase of gluconic acid.
One of the key reasons for using these types of sensor in an intravascular environment, rather than subcutaneously or in other bodily environments, is the need to provide closed-loop control for diabetic patients. This would provide insulin delivery based on the patient's actual glucose measurements, as opposed to providing insulin based on some inexact approximation of the patient's glucose levels. This would of great benefit to diabetic patients. There is a widely recognized time delay between glucose changes in venous blood, and subcutaneous glucose changes. This time delay can range from just a few minutes, to up to 30 min. However, the mathematical algorithm used to couple the glucose signal to the insulin delivery system cannot tolerate a very long time delay. In fact, two authors have presented data which suggested that 10 minutes is the maximum delay which can be tolerated in closed-loop insulin delivery systems (Parker R S, Doyle F, et al., “A Model-Based Algorithm for Blood Glucose Control in Type I Diabetic Patients” IEEE Trans. Biomed. Engr. 46(2):148-157 (1999), and Gough D et al, “Frequency Characterization of Blood Glucose Dynamics” Ann Biomed Engr 31:91-97 (2003).) Longer time delays can cause the controller to become unstable, potentially creating life-threatening issues for the patient, such as delivery of extra insulin when blood glucose levels are falling rapidly.
Despite the foregoing and other efforts in the art, a suitable continuous in dwelling glucose sensor has not yet been developed.
A critical factor in the design of an implanted sensor is the anatomical site in which it is implanted. A few investigators have developed monitoring systems, which can be placed within the vascular system. Armour et al. (“Application of Chronic Intravascular Blood Glucose Sensor in Dogs”, Diabetes 39:1519-26 (1990)) implanted a sensor into the superior vena cava of six dogs for a period of up to 15 weeks with relative success. However, due to the risks of thrombosis and embolization, the majority of investigators have focused on subcutaneous implantation.
A major drawback to subcutaneous implantation is the body's defense against foreign objects: the “foreign-body response”. In this host response, if an object cannot be removed by the inflammatory response, foreign-body giant cells will form a “wall” around the object, which is subsequently followed by the formation of a fibrous capsule. If the object is a blood glucose sensor, it will no longer be in intimate contact with body fluids, and the signal will drift and stability will be lost. There are numerous reports of sensor stability being lost in about a week (Wilson, G. S., et al., “Progress Towards The Development Of An Implantable Sensor For Glucose”, Clin. Chem. 1992 38:1613-7, and Kerner, et al., “A Potentially Implantable Enzyme Electrode For Amperometric Measurement Of Glucose”, Horm. Metab. Res. Suppl. Ser. 20: 8-13 (1988)). Updike et al. (Updike, Stuart J., et al., “Enzymatic Glucose Sensors: Improved Long-Term Performance In Vitro And In Vivo”, ASAIO J., 40: 157-163 (1994)) reported on the subcutaneous implantation of a sensor which was stable for up to 12 weeks, however, this evaluation was only performed in three animals.
Recent clinical studies have also demonstrated that implantable insulin pumps are feasible for implantation for over one year (Jaremko, J. et al., “Advances Towards the Implantable Artificial Pancreas for Treatment of Diabetes,” Diabetes Care, 21(3): 444-450 (1998)). The research was inspired by the goal of the development of the artificial pancreas, and promising initial clinical trials using implantable insulin pumps. At this point in time, development of implantable insulin pumps is at a very advanced stage, with units being implanted for over 2 years in canines (Scavani et al., “Long-Term Implantation Of A New Programmable Implantable Insulin Pump,” Artif. Organs, 16: 518-22 (1992)) and in 25 patients for up to 3 years (Waxman, et al., “Implantable Programmable Insulin Pumps For The Treatment Of Diabetes”, Arch. Surg., 127: 1032-37 (1992)).
A number of wearable insulin pumps are described by Irsigler et al. (“Controlled Drug Delivery In The Treatment Of Diabetes Mellitus,” Crit. Rev. Ther. Drug Carrier Syst., 1(3): 189-280 (1985)). Thus, it should be relatively straightforward to couple a long-term implantable glucose sensor as described in this disclosure, to an insulin pump to optimize glycemic control for the patient.
In another aspect of this invention, it is possible to apply the principles discussed above to the direct, continuous monitoring of arterial blood gases (ABG). Arterial blood gas values such as pO2, pCO2, and pH are the most frequently ordered laboratory examinations in the intensive care setting and the operating room (C. K. Mahutte, “Continuous intra-arterial blood gas monitoring,” Intensive Care Med (1994) 20:85-86). In the intensive care unit (ICU), ABG is typically monitored once a day, and additional measurements are only made once the patient has experienced a deleterious event. Limited additional sampling is performed at the discretion of a physician or nurse. There can be a significant time delay between the time the tests are ordered, and the results are returned (E. E. Roupie, “Equipment Review: Continuous assessment of arterial blood gases,” Crit. Care 1997 1(1):11-14).
In the case of continuous monitoring, significant changes in ABG values or trends would cause a rapid, therapeutic response on the part of the physician, so potentially catastrophic events could be avoided. Continuous, non-invasive monitoring techniques such as pulse oximetry and continuous capnography have been introduced for this reason. Unfortunately, these devices are not always accurate in cases such as shock, hypothermia, or during the use of vasopressors. Further, pulse-oximetry does not measure oxygen tension.
A number of attempts have been made to develop improved arterial blood gas monitors. There are two basic types of arterial blood gas monitors. In the first type, termed extra-arterial blood gas (EABG) monitors, the patient's blood gas values are measured from a sample in the arterial catheter. This can significantly reduce the time delay in obtaining results, as compared to sending the sample to a laboratory. However, it is an on-demand system, not a continuous one, so the frequency of sampling is once again dependent upon the physician or nurse.
The second type of device, known as an intra-arterial blood gas (IABG) monitor, is inserted directly into the arterial blood. However, the consistency and reliability of these IABG monitors have not been clinically acceptable because of problems associated with the intra-arterial environment. (C. K. Mahutte, “On-line Arterial Blood Gas Analysis with Optodes: Current Status,” Clin. Biochem 1998; 31:119-130).
In another aspect of this invention, implantable sensors capable of monitoring hemodynamic conditions over extended periods of time may be useful in patients with heart failure. Such measurements have traditionally been restricted to cardiac catheterization laboratories and intensive care units (ICU's). Measurements cannot be made easily in an ambulatory setting, or under conditions of cardiac loading, such as exercise. Ohlsson et al (Ohlsson A, et al., “Continuous ambulatory monitoring of absolute right ventricular pressure and mixed venous oxygen saturation in patients with heart failure using an implantable hemodynamic monitor: Results of a one year multicentre feasibility study,” Eur Heart J 22:942-954 (2001)) describe implantation of a pressure sensor and an oxygen sensor, the IHM-1, Model 10040 (Medtronic, Inc.). However, in this study, 12 out of 21 oxygen sensors failed within the first 6 months of implantation. A fibrinous coating covered one of the sensors, and it was believed that this was responsible for sensor failure. In addition, surgical implantation of this pacemaker-style device requires a 2-3 hour procedure in the operating room, which is more costly than a catheterization procedure to insert a pressure sensor. Data from this study, however, points to the importance of long-term, continuous pressure monitoring, as opposed to one-time measurements using standard cardiac catheterization techniques.
In another aspect of this invention, the sensor may be used for measuring flow rates within a vessel. Congestive heart failure affects more than 5 million persons in the United States, and the rate is increasing as people age and more of them survive heart attacks. Finding the right treatment often involves a trial-and-error process, with the physician trying different combinations of drugs and different dosages to produce the best results. That means repeated blood pressure and flow measurements and invasive cardiac catheterization procedures to measure the effects on the heart.
In U.S. Pat. No. 6,053,873 issued Apr. 25, 2000 to Govari et al, a method of monitoring flow rates within a stent is described. In particular, the use of ultrasonic sensors to monitor blood flowrates is described. However, due to the change in composition and thickness of the biological material present on the sensor surface, the boundary conditions at the sensor may change over time. Initially, the sensor surface, which may be calibrated against blood, may become covered with a layer of thrombus over time. Subsequently, the thrombus will transform into fibrous tissue, affecting the absorption of the ultrasound signal. Blood may have an ultrasound absorption level similar to water, of around 0.002 dB/MHz cm, while the fibrous tissue layer may have an absorption level similar to muscle, of around 2 dB/MHz cm. Thus, such a sensor may require periodic re-calibration. Unfortunately, the re-calibration process requires an invasive cardiac catheterization procedure, in which flow rates are determined using ultrasound, thermodilution, or the Fick method. In order to avoid additional interventional procedures, the fouling-resistant approach described herein may be valuable.
In U.S. Pat. No. 6,309,350 to VanTassel, et al., issued Oct. 30, 2001, a sensor is described which is anchored in the wall of the heart, or in a blood vessel. However, no mention is made of the difficulty with sensor fouling due to thrombus formation, or how to compensate for that. Pressure sensors would require recalibration as the softer thrombus transforms into neointimal tissue, or as the thickness of the encapsulating tissue layer changes. While the device could be initially calibrated during implantation, it would need to be recalibrated over time. Because recalibration requires an invasive cardiac catheterization procedure, it would be desirable to avoid this if at all possible.
In addition, VanTassel et al. suggest the use of standard thermodilution methods to determine flow rates. Thermodilution is normally performed by injection of either room temperature or iced saline through a catheter. However, it is desirable to avoid re-interventions if possible. Therefore, standard thermodilution methods are less desirable than the current invention. In addition, it is impractical to use the sensor in the standard method, i.e., by locally cooling the blood, because chilling units are impractically large for implantation as a sensor. Further, thermodilution methods would be inaccurate if the thermocouple or sensing element was covered by a relatively thick layer of cells. This would be especially true if only a small temperature rise were introduced into a large volume of flowing blood, such as in the pulmonary artery. Larger temperature rises (>2.5° C.) would cause local tissue damage.
Stroke, also known as cerebral infarction, is the third most common cause of death in the US and Europe. There are approximately 1 million acute ischemic strokes in Europe annually. Stroke survivors are often significantly disabled, and must undergo extensive rehabilitation. The estimated short term costs of a stroke are $13,649 per patient, and the long term costs are $45,893 for a minor stroke and $124,564 for a major stroke (Caro, J J, et al., “Stroke Treatment Economic Model (STEM): Predicting Long-term Costs from Functional Status”, Stroke 1999; 30:2574-2579). The direct and indirect economic impact of stroke in the US is estimated to be $43.3 billion (US). According to the National Institutes of Health recombinant tissue-type plasminogen activator trial, if stroke could be treated in the first hours after the stroke occurs, the damage can be minimized (National Institute of Neurological Disorders and Stroke. rt-PA Stroke Study Group. “Tissue Plasminogen Activator for acute ischemic stroke.” N Engl J Med 1995; 333:1581-1587). However, if treatment is delayed, then the damage becomes worse. Unfortunately, symptoms of stroke are not widely recognized, and significant delays in treatment are introduced by the failure to recognize the problem.
Risk factors such as high blood pressure, cholesterol, smoking, obesity, and diabetes increase the chances that an individual will suffer a stroke. Patients with significantly increased risks of stroke include patients with atrial fibrillation, coronary heart disease, asymptomatic carotid stenosis, previous stroke, or transient ischemic attack (TIA). In the latter subsets, approximately 25% of stroke survivors experience a recurrent stroke within five years, and among patients with TIA, the risk for stroke is 5% within 48 hours, 12% within 1 year, and up to 30% within 5 years.
In addition, each year, a significant number of surgical procedures are performed in which the patient is exposed to the risk of stroke in a hospital setting. For example, according to the American Heart Association, there are approximately 500,000 coronary artery bypass grafting (CABG) procedures performed annually. The risk of stroke during or shortly after these procedures is about 5%. Carotid endarterectomy is another common surgical procedure which is intended to reduce the long-term risk of stroke, but still carries significant short-term stroke risk of about 5-10%. In-hospital monitoring would be suitable for patients undergoing cardiac surgery, large vessel surgery, intra-cranial surgery, or other procedures that may place them at substantially increased risk of stroke either during the procedure, or during the post-operative recovery period.
The risk of stroke either during or following surgery is significantly increased for a period of approximately one month (David J. Blacker, Kelly D. Flemming, Michael J. Link, Robert D. Brown, Jr. “The Preoperative Cerebrovascular Consultation: Common Cerebrovascular Questions Before General or Cardiac Surgery” Mayo Clin Proc. 2004; 79:223-229, and PCA Kam, R M Calcroft, “Perioperative stroke in General Surgical Patients,” Anesthesia 1997; 52:879-883). Strokes which occur in-hospital often go untreated for significant periods of time, for a variety of reasons including misdiagnosis, misinterpretation of stroke symptoms as due to other factors, and the referring team not being familiar with acute stroke assessment (David J Blacker, “In-hospital stroke” Lancet Neurol 2003; 2: 741-46). Since the cost of stroke is so significant both in financial terms and in the loss of quality of life, it would be desirable to alert the physician to the occurrence of stroke as quickly as possible, so that permanent damage may be prevented or reduced. Therefore, it would be desirable to have a stroke-monitoring device that can be used for this period of time. Further, it would be desirable to have a portable monitoring system, which the patient can use whether in the hospital, or following discharge.
Current stroke monitoring methods provide responses that are generally too slow to allow tissue rescue to be performed by the physician. Currently, one method of perioperative stroke monitoring, somatostatic excitatory potentials (SSEP) is rather slow, and often does not provide physicians with sufficient warning to alter patient outcomes. In instances where the motor pathways and sensory pathways are not in close proximity, significant loss of motor function can occur without detection by SSEP. Another method, which may be statistically correlated with stroke, is transcranial Doppler (TCD). Using this method, both reduced flow rates and emboli may be detected. However, this does not detect stroke directly. Since there are usually a number of embolic events during a surgical procedure, it would be desirable to alert the physician only when these emboli cause a stroke.
In the acute phase of stroke, or cerebral infarction, a great deal of experimental data suggests that free radicals, including superoxide, hydroxy radical, and nitric oxide (NO) are one of the most important factors causing brain damage. J. Rodrigo, D, et al (Histol. Histopathol. 17, 973-1003 (2002)) have observed that most of the morphological and molecular changes associated with ischemic damage were prevented by treatment with inhibitors of NO production. In addition, there is a significant increase in the concentration of NO following cerebral ischemia. There is also significant NO release in hemorrhagic stroke. (Chen H H, et al, “Low cholesterol in erythrocyte membranes and high lipoperoxides in erythrocytes are the potential risk factors for cerebral hemorrhagic stroke in human” Biomed Environ Sci. 2001 September; 14(3):189-98). Therefore, the detection of NO in the acute phase of ischemic or hemorrhagic stroke would provide an early detection method for stroke, allowing treatment to be performed as promptly as possible by the physician or health-care provider.
Animal studies have demonstrated an increase in nitric oxide in brain tissue shortly after ligation of the middle cerebral artery (Lin, S Z, et al., “Ketamine Antagonizes Nitric Oxide Release From Cerebral Cortex after Middle Cerebral Artery Ligation in Rats”, Stroke 1996; 27:747-752). It has further been demonstrated that there is a significant increase in levels of nitrite and nitrate ions in the jugular vein of a rat, immediately following induction of ischemic stroke (Suzuki, M, et al., Brain Research 951 (2002) 113-120). It has also been demonstrated that the level of these metabolites are significantly higher in cerebrospinal fluid (CSF) at the time stroke patients are admitted to the hospital, which is often many hours after the stroke (Castillo, J. et al., “Nitric Oxide-Related Brain Damage in Acute Ischemic Stroke”, Stroke 2000; 31:852-857). Increased nitrate and nitrite levels correlate with increased infarct volume and poorer neurological outcomes.
Nitric oxide sensors have been coupled with catheters and used in vivo in a canine aorta in order to monitor sub-nanomolar changes in NO release in response to acetylcholine stimuli (Seiichi Mochizuki, Takehiro Miyasaka, Masami Goto, Yasuo Ogasawara, Toyotaka Yada, Maki Akiyama, Yoji Neishi, Tomohiko Toyoda, Junko Tomita, Yuji Koyama, Katsuhiko Tsujioka, Fumihiko Kajiya, Takashi Akasaka, and Kiyoshi Yoshida “Measurement of acetylcholine-induced endothelium-derived nitric oxide in aorta using a newly developed catheter-type nitric oxide sensor” Biochem Biophys Res Comm 306 (2003) 505-508). Unfortunately, there are a number of problems associated with simple catheter-sensor combinations, as discussed for the intra-arterial blood gas sensors that have been investigated clinically (C. K. Mahutte, “On-line Arterial Blood Gas Analysis with Optodes: Current Status,” Clin. Biochem 1998; 31:119-130). These problems include thrombus formation at the tip, measurement of analyte concentrations in the vascular wall instead of the blood stream, vessel compression due to patient movement, and vasospasm. In addition, signal variation due to catheter movement is also a potential problem. All of these issues are addressed in the present invention.
Other challenges with the measurement of nitric oxide in vivo include its low concentration (in the nanomolar range), and its short half-life (3-5 sec). (DS Bredt, “Endogenous Nitric Oxide Synthesis: Biological Functions and Pathophysiology”, Free Rad. Res. 31(6):577-596 (1999).) Nitric oxide electrodes are also sensitive to variations in flow rate and temperature, further complicating the measurement of NO in blood.
A potential solution to this challenge is to monitor the metabolic products of NO, which include nitrite and nitrate. However, nitrite (NO2), one of the metabolic products of nitric oxide, also has a relatively short half-life of about 10 minutes in blood. Therefore, even analysis of nitrite levels in blood using HPLC requires that chemical stabilizers be added to prevent the decomposition of nitrite prior to analysis. The available methods for measurement of nitrite levels in blood and plasma based on mass spectroscopy and chemiluminescence require several chemical and physical preparatory steps and the use of standard curves. The samples have to be grouped for a proper assessment and the nitrite levels needs to be determined in a delayed fashion, with the most reliable results reported after finishing the entire sample collection and analysis. Nitrite and nitrate have widely varying background levels, which are influenced by a number of factors including diet. There are currently no commercially available NO2 sensors that can directly measure NO2 in the physiological concentration range in blood or in electrolyte solutions (such as ultrafiltrate solutions) obtained from blood (0.5 to 2 μM NO2).
Variable background contamination of water and laboratory supplies are known to be issues in the measurement of nitrites and nitrates. “Considering the frequent need to measure NO in the nanomolar range when dealing with biological samples, the nitrite contamination found in tap water is highly problematic. Further, because of the ubiquitous nature of nitrite, the use of substantially nitrite-free water is imperative not only when diluting standards and samples, but also when rinsing glassware and syringes. The importance of using fresh filtered or glass-bottled water is underscored by the fact that NO from the air can react with oxygen in water to form nitrite. The use of substantially nitrite-free water in solutions such as the preservation solution and to dissolve EDTA is also important in order to eliminate nitrite contamination in these solutions.” (Peter H. MacArthur, Sruti Shiva, Mark T. Gladwin, “Measurement of circulating nitrite and S-nitrosothiols by reductive chemiluminescence”, Journal of Chromatography B, 851 (2007) 93-105.) “Nitrite, the oxidative breakdown product of NO and a ubiquitous trace contaminant of chemicals and laboratory glassware, is a well-known nuisance in NO analytical chemistry.” (Xunde Wang, Nathan S. Bryan, Peter H. MacArthur, Juan Rodriguez, Mark T. Gladwin, and Martin Feelisch, “Measurement of Nitric Oxide Levels in the Red Cell Validation Of Tri-Iodide-Based Chemiluminescence With Acid-Sulfanilamide Pretreatment” J Biological Chemistry Vol. 281, No. 37, pp. 26994-27002 (2006).)
In particular, ultrafiltration membranes can be a source of nitrite and nitrate contamination, and therefore their use is challenging. “Use of ultrafiltration units to remove protein from plasma is a cause of heavy contamination that persists to a certain degree even after several washes with pure water.” (Takaharu Ishibashi, Junko Yoshida, and Matomo Nishio, “New Methods to Evaluate Endothelial Function: A Search for a Marker of Nitric Oxide (NO) In Vivo: Re-evaluation of NOx in Plasma and Red Blood Cells and a Trial to Detect Nitrosothiols” J Pharmacol Sci 93, 409-416 (2003).) In addition, the time delay imposed by ultrafiltration of samples means that measurement of nitric oxide itself from a sample is unlikely.
In addition to monitoring for cerebral ischemia, monitoring NO2 concentrations may also have applications in ischemia in general, including cardiac ischemia, in septic shock (660,000 patients/year) (Kuhl S J, Rosen H. “Nitric oxide and septic shock—from bench to bedside.” West J Med 1998; 168:1 76-181; and Greg S. Martin, M.D., David M. Mannino, M.D., Stephanie Eaton, M.D., and Marc Moss, M.D. “The Epidemiology of Sepsis in the United States from 1979 through 2000”, NEJM 348:1546-1554 (2003).), traumatic brain injury (280,000 patients/year) (Hlatky, R et al, Role of Nitric Oxide in Cerebral Blood Flow Abnormalities After Traumatic Brain Injury, J Cereb Blood Flow Metab 23:582-88 (2003).) subarachnoid hemorrhage (SAH) (21,000 patients/year) (R Loch Macdonald, Ryszard M Pluta and John H Zhang, Nature Clinical Practice Neurology May 2007 3(5) 256-63; and Writing Group Members, Thomas Thom, et al., “Heart Disease and Stroke Statistics—2006 Update. A Report From the American Heart Association Statistics Committee and Stroke Statistics Subcommittee”, Circulation February 2006; 113:e85-e151) asthma, and pain management (A. Koch, K. Zacharowski, O. Boehm, M. Stevens, P. Lipfert, H-J. von Giesen, A. Wolf and R. Freynhagen, “Nitric oxide and pro-inflammatory cytokines correlate with pain intensity in chronic pain patients” Inflamm. res. 56 (2007) 32-37), erectile dysfunction, pulmonary hypertension, and other pathologic conditions as known to one of skill in the art.
Notwithstanding the efforts in the prior art, however, there remains a need for intravascular sensors for implantation or insertion in a blood vessel, which can provide useful blood glucose or other physico-chemical readings for an extended period of time, without material interference from thrombus formation, embolization, or other foreign body response. Preferably, the sensor is capable of continuous or near continuous monitoring, and driving an implantable insulin pump and/or making blood glucose or other data available to the patient or medical personnel.